Surgical “vaporization” of tissue, or “ablation”, is simple as a physical concept but is chemically quite complex. While the dominant constituent of most human tissues is water and water does truly vaporize upon heating, the organic components (proteins, fats, nucleic acids, sugars) generally thermally decompose, or burn, before becoming vapor. Heat must be produced in tissue to vaporize water and decompose and then vaporize organic tissue components, but laser energy is not necessarily the source of all of the heat doing work in surgery; much of the energy involved derives from exothermic processes within the tissues that are stimulated by the laser energy, much as a match to a candle.
In soft tissues, laser radiation is converted to vibrational energy at some point during a cascade of simultaneous chemical reactions that take place. Even where chromophores strongly absorb the surgical laser wavelength, the local energy density is far below that required for photodissociation: target compounds absorb light and convert photonic energy into vibrational energy through electronic and phononic interactions. Some of the vibrational energy in chromophore(s) is also conducted (as heat) to adjacent compounds that absorb the incident laser radiation less efficiently than the primary chromophore(s). Where secondary species are in close proximity to primary chromophores, or are relatively volatile (e.g. water), they may decompose or vaporize rapidly. Where the thermal conduction path is longer, organic tissue components may be thermally damaged but not immediately removed (coagulation), which may increase post-operative complications and can even result in death; tissue coagulation is a complication of laser surgery in soft tissues that may result in hyponatremia and is to be avoided.
The efficiency of chromophore absorption, and the concentration of the chromophore, affects the depth of penetration of the incident laser light: given equivalent power density, the more concentrated the chromophore and more efficient the absorption, the shallower the light will penetrate a given tissue. Shallow penetration necessarily concentrates laser radiation in smaller tissue volumes and the underlying coagulation zones are concomitantly thin. At deep depths of penetration, significant volumes of tissue underlying the target tissue are irradiated at power densities insufficient for vaporization and are thus coagulated: a shallow depth of penetration exposes smaller underlying volumes of tissue to sub-therapeutic energy densities. Similarly, as laser power applied to target tissues is increased, the depth of penetration of damaging, sub-therapeutic energy also increases and more non-target tissue volume is coagulated. A laser wavelength that offers extremely high absorptivity in target tissues may be used at much higher average power with the same or less collateral tissue damage (coagulation) than a laser wavelength that is less efficiently absorbed.
For most surgical lasers, the target chromophore(s) is but a fraction of the tissue mass; where water is the chromophore it is a considerable fraction, but in that water is also a ubiquitous constituent of tissue, it is far from a tissue-specific target; labeling water a ‘chromophore’, in fact, is dubious given its broad absorption spectrum and lack of tissue specificity. Non-targeted tissue components absorb laser light less efficiently and less efficient absorption results in scattering of the incident beam. This more or less “non-specific absorption” also contributes to sub-optimal tissue heating and collateral tissue damage (coagulation) of adjacent and underlying tissues through thermal contact. Inefficient laser absorption by tissue breakdown products also leads to “charring”: the carbonization of organic tissue components that is familiar to any backyard grilling enthusiast. In surgery, charred tissue scatters laser energy quite strongly and under mechanisms (and affects) that differ with the size (mean diameter) of the carbon particles and particle aggregates, adding to collateral tissue damage and rendering the final therapeutic outcome far less predictable.
Whether water and/or hemoglobin strongly absorb the laser energy has evolved as a discriminating characteristic for laser manufacturers in marketing campaigns (as well as in patent claims) and has spawned a new term—“photoselectivity”—a word that is used to infer unique precision in targeting a “specific” tissue component (hemoglobin or H2O). Lasers that target water as the primary chromophore (e.g. NIR diodes, holmium, thulium, erbium) are deemed ‘photoselective’ for water in that most of the other tissue components do not absorb these lasers' wavelengths strongly (but all tissues contain water). Lasers said to be photoselective for hemoglobin, e.g. Greenlight HPS, are only photoselective for as long as hemoglobin exists in the target tissue—on the order of femtoseconds post initial exposure—after which these laser wavelengths affect tissues much the same as any other laser wavelength.
Concomitant with the desired laser-tissue interactions during laser surgery, simultaneous and undesirable processes occur that cause damage to non-target tissues and to the surgical fiber optic probe. Where the surgical goal is to affect vaporization of soft tissues, coagulation of adjacent and underlying tissues occurs to a greater or lesser extend depending upon the characteristics of the laser light used, e.g. wavelength, average power density, beam profile, instantaneous energy density, variable surgical expertise and the light absorption characteristics of the target tissue, e.g. concentration of chromophores, scattering centers, tissue density, thermal conduction. Coagulated and chemically altered tissues adhere to the surgical probe, absorbing energy intended for target tissues and heating the probe instead, often to temperatures that severely damage the probe and nearby tissues, i.e. up to temperatures resulting in incandescence and even melting of the fused silica probe tip.
At any power level of current visible and near infrared lasers used in ablating tissue, heat ultimately does the bulk of the work and causes the bulk of the problems. Ideally, this heat is confined to as small a tissue volume as possible during application of energy such that affected tissues are ablated and not merely coagulated, but inefficient delivery of even highly absorbed radiation by damaged surgical probes results in undesirable outcomes, e.g., generalized heating, that may cascade into even more undesirable outcomes.
Non-radiative relaxation heats the system (the tissue microenvironment) where radiative relaxation reintroduces photonic energy to the microenvironment. The complexity of the resulting microenvironment therefore increases as new vibrational states are created and photons of different energy are emitted. More complexity is introduced as chemical bonds break and form, creating new molecules that absorb and radiate energy differently than the untreated tissue. The intense laser irradiance in surgery brings forth additional phenomena such as non-linear absorption (multi-photon absorption) and saturable adsorption (loss of absorptivity due to saturation of excited states).
Char is formed from the carbonization of tissue and is thermally driven. Development of char on tissue during laser surgery is advocated for strongly absorbing visible and near IR wavelengths, purportedly improving surgical efficiency but it also increases coagulation and results in slower localized healing. In particular, carbon particles with a high refractive index scatter incident laser radiation to create higher collateral tissue damage. Two models of light scattering by carbon particles must be considered in describing the adverse scattering: Rayleigh and Mie scattering.
The Rayleigh scattering model applies to particles much smaller than the wavelength scattered. The scatter intensity increases as the ratio of particle size to wavelength increases, up to particle mean diameters of approximately 10% of the wavelength (after which the model breaks down). Rayleigh scatter is equally intense in forward and reverse directions but the overall intensity is small relative to the incident energy. As carbon particles grow in diameter past approximately 10% of the incident wavelength, Mie scattering comes to dominate. Mie scatter is very intense in comparison to Rayleigh scatter and it is proportional to the square of the particles' cross-sectional area—increasing exponentially with particle size—and forward scatter angles are favored. Accordingly, a small amount of char may improve tissue removal rates through Rayleigh scattering, initially, but as the particles grow to favor Mie scatter, collateral tissue damage increases much more rapidly than ablation efficiency; even a little char is undesirable.
A common chromophore ‘targeted’ by surgical lasers is oxyhemoglobin (HbO2): the most abundant, true chromophore endogenous to most tissue. Surgical laser companies use the strong absorption of HbO2 in rationalization of high performance for any laser wavelength that falls at, or relatively close to, even minor absorptivity peaks. In situ, laser induced changes in oxyhemoglobin include deoxygenation, loss of the heme-porphyrins and formation of free heme, denaturalization of the protein, thermal decomposition, etc. This laser induced chemistry proceeds very quickly and produces a blizzard of complex chemical reactions including coupling of N-terminal amino acids and reducing sugars, e.g. deoxyribose, Amadori and Maillard reactions, and ultimately carbonization.
The most successful soft tissue surgical laser to date is the CO2 laser. Water, HbO2 and most other tissue constituents strongly absorb the CO2 laser wavelength of 10.6 μm (10,600 nm) but flexible waveguides suitable for endosurgery applications, such as vaporization of hyperplastic prostate tissue, have proven elusive, limiting CO2 laser applications in surgery primarily to external tissues, e.g. plastic surgery. Early laser treatment of benign prostatic hyperplasia (BPH) used Nd:YAG lasers and frequency doubled Nd:YAGs (aka KTP lasers at 532 nm) and silica optical fiber probes but owing to relatively deep penetration within the target tissue, the 1064 nm wavelength coagulated far more tissue than it vaporized and post-operative complication rates were high.
Beginning in the late 1990s, Ho:YAG lasers largely replaced Nd:YAG lasers for BPH surgery because the 2140 nm wavelength penetrates tissue much less than 1064 nm due to a strong absorption by water, but the pulse output of holmium lasers leaves ragged tissue edges while the wavelength and high peak pulse energies stress fiber optic surgical probes while strong absorption by water—also used in sterile irrigation of the surgical field—causes boiling at probe tips and strongly attenuates the laser energy unless the probe is maintained in close contact with the target tissue (where tissue contact is damaging to the surgical probe).
Frequency doubled Nd:YAG/KTP lasers reached 532 nm output powers capable of reasonably efficient vaporization by the early 2000s. Prostate tissue proved to absorb the “greenlight” fairly well while the laser energy passed through water unaffected, but the single HbO2 target chromophore was found to “bleach” rapidly, greatly reducing vaporization efficiency as surgery progressed. Tm:YAG lasers appeared in the mid-2000s, offering a 2000 nm wavelength similar to the Ho:YAG, but continuous rather than pulsed with water strongly absorbing the energy: surgical probe to tissue separation issues remain similar to Ho:YAG.
High power diode lasers were presented as lower cost alternatives to the solid state lasers starting at 880 nm and progressing to 980 nm and 1470 nm while 532 nm laser output was increased twice in the next decade, from 80 W to 120 W and finally to 180 W. All of the infrared lasers then joined in the power race with a 980 nm/1470 nm dual wavelength diode laser producing 200 W of combined power (or 160 W at 980 nm and 40 W at 1470 nm, individually or blended, said to offer a continuum of tissue response from vaporization and coagulation), a 250 W diode laser at 980 nm, a Ho:YAG offering 120 W average power (pulsed) and a 200 W Tm:YAG.
Alone, higher power lasers present diminishing returns, particularly for wavelengths that target HbO2, the higher power caused photobleaching of the target tissue and significantly reduced hemostasis (i.e., more bleeding with more power). While high power lasers cause deeper coagulation, they increase expenditures due to consumption of the surgical fibers used to deliver the laser energy to, for example, the prostate. Notably, higher laser powers consume fibers more quickly or require larger, more expensive optical fiber, e.g. a 532 nm laser at 80 W or 120 W requires a 600 μm core fiber whereas at 180 W requires a liquid-cooled 750 μm core fiber.